Defibrillation waveforms are typically defined by the energy expended during the delivery of the waveform. The parameters defining the energy are voltage, current and time. Generally, most defibrillator devices use a preset voltage, and assume the impedance of the heart to be about 50.OMEGA.. Thus, only the time parameter is a variable. This means that energy delivery is, at best, an approximation based on a range of preset values. The operational specifications for defibrillation waveforms in accordance with the standards of the Association for the Advancement of Medical Instrumentation (AAMI) is indicated to be at a 40% level of accuracy across impedance ranges.
Prior art devices use a single measurement of current at the beginning of a waveform to calculate the resistance. The duration or time is then adjusted to construct the correct energy during the delivery of the defibrillation pulse. One of the limitations of this approach is that it assumes the resistance remains constant throughout the defibrillation pulse delivery. Studies regarding capacitative discharge waveforms demonstrate that electrical impedance increases as voltage decreases, and that this relationship is not simply due to the Ohm's Law. For the stimulation of excitable biological tissues, current decreases at a faster rate than does voltage. This non-linear relationship between impedance and voltage is largely due to the electrode-tissue interface. (Low Voltage Shocks have a Significantly Higher Tilt of the Internal Electric Field Than Do High Voltage Shocks, by James E. Brewer, et al, Pacing and Clinical Electrophysiology, Volume 18, No. 1, January 1995).
The assumption that resistance remains constant is particularly erroneous in light of the transchest/transthoracic application shock delivery involving external defibrillators (EDs). The transchest discharge of shock pulses involves chest resistances which include chest wall resistance, lung series resistance, lung parallel resistance, thoracic cage resistance, in addition to resistance of heart and heart cell membrane. Further, external defibrillators are generally used on random patients under emergency situations. Thus patient variability increases the possible variance of resistance across patients. Realizing this, the assumptions and design parameters of the prior art are inadequate to provide a robust and reliable defibrillation pulse delivery suited for implementation in ED devices.
The shape of a defibrillation waveform is a critical element in the successful delivery of a defibrillation waveform. Specifically, prior art defibrillation pulses for monophasic or biphasic waveforms include preset parameters which are used to structure the waveform. For biphasic waveforms, these parameters include an initial voltage of .phi..sub.1 and its duration or tilt, and an initial voltage of .phi..sub.2 and its duration or tilt. Additionally, the interphase duration between .phi..sub.1 and .phi..sub.2 is set to account for interphase delay required to switch from one phase to the other.
Prior art waveform generation assumptions and calculations contain several limitations. Specifically, the waveforms are not congruent with the myocardial cell response. The theory of myocardial cell response is based on the observation that .phi..sub.1 defibrillates the heart and for biphasic waveforms .phi..sub.2 performs a stabilizing action that keeps the heart from refibrillating by canceling out (burping) any residual charges in the myocardial cells. For monophasic waveforms, these parameters include the initial voltage and the pulse duration or tilt. Further, the efficacy and advantages of monophasic and biphasic waveform pulses is significantly enhanced if the phases are shaped to simulate the myocardial cell response waveform. For biphasic waveforms, the congruence in simulation between the waveform and the myocardial cell response requires that the residual charges be removed immediately after the delivery of .phi..sub.1. This implies that there should be no interphase delay between the phases. Moreover, the dynamically changing parameters such as the resistance and the myocardial cell response require that the waveform tilt equation and the duration of the phases must take these variables into account.
Accordingly, there is a need to provide a method and device to enable an autonomous and accurate delivery of a monophasic or biphasic defibrillation waveform which is compatible with the dynamic myocardial cell response in a variable resistance environment.